Fluorapatite‐Coated Percutaneous Osseointegrated Prostheses Limit Epidermal Downgrowth and Promote Periprosthetic Healing in a Weight‐Bearing Sheep Model—A Preliminary Study

ABSTRACT

Percutaneous osseointegrated (OI) devices offer an advanced alternative to socket prosthetic suspension systems for amputees but can face clinical limitations due to complications at the skin‐implant interface. Titanium (Ti) and its alloys, though mechanically suitable and promoting osseointegration, are unable to support epidermal cell adhesion, leading to chronic wounds, sinus tract formation, and infections. To improve epidermal tissue integration, this study explored the use of fluorapatite (FA)—a natural tooth mineral known to allow epithelial adhesion through hemidesmosomes. It was hypothesized that FA‐coated Ti surfaces would enhance epidermal cell attachment, promote wound healing, and prevent epithelial downgrowth. To test this idea, FA‐coated Ti devices were implanted into the transected fused 3–4 metacarpals of five sheep and evaluated at 12 weeks post‐implantation. The results revealed that FA‐coatings significantly improved healing outcomes and limited epithelial migration along the implant surface. These findings support FA as a promising strategy for promoting a biological seal around percutaneous implants.

1. Introduction

Percutaneous osseointegrated (OI) prosthetic devices have proved to be a successful alternative to traditional socket‐suspended prostheses. This approach involves inserting an endoprosthesis into the medullary canal of the transected bone, with an extension which then exits the skin barrier and provides a platform for affixing an exoprosthesis [1, 2, 3, 4, 5]. Titanium OI devices have demonstrated significant utility in various medical applications [6, 7, 8, 9, 10, 11, 12, 13], offering superior structural support by providing a stable OI skeletal docking platform for attaching artificial limbs. They have fulfilled their promise of revolutionizing amputee rehabilitation by obviating the historical challenges of socket‐suspended prosthetic ambulation [3, 14, 15, 16]. The direct skeletal attachment allows patients to perceive ground strikes and terrain changes when walking, allowing them to perceive the prostheses as part of their own body; this has been termed osseoperception. [17, 18] The complications inherent in socket‐suspension systems, such as pressure sores, discomfort during sitting, tissue necrosis, and inadequate load transfer to the residual bone are no longer an issue [19]. Limited prosthetic wear times, often resulting in the abandonment of the artificial limb and leaving amputees wheelchair‐bound are eliminated [5, 20, 21, 22, 23, 24, 25]. While skeletally attached artificial limbs are functionally superior, the skin‐to‐implant interface at the device exit site remains a significant weak link [26, 27, 28]. Since the implant disrupts the protective skin barrier at the stoma, healing responses are complex, prolonged and persistent [29]. Suboptimal healing cascades at the skin/implant interface can lead to the perpetual presence of inflammation, accompanied by the formation of granulation tissue (GT).

Because of bone's well‐known ability to integrate with titanium‐based implants (i.e., osseointegration), medical‐grade titanium alloys are the preferred materials for fabricating percutaneous OI prosthetic devices. Although foreign body responses (FBRs) in the exiting soft‐tissue sites are expected to occur following implantation, they have been historically less well studied in this device type [30, 31]. In short, the elicited FBR is a healing process that attempts to eliminate or isolate the foreign body (i.e., implant) from the soft tissue through inflammation (phagocytosis) and/or biological encapsulation [29]. Simultaneously, at the distal implant exit site (stoma), the presence of a skin‐protruding device and the inability of epithelial cells to integrate with the device's surface hinder the completion of the re‐epithelialization phase of the wound healing cascade. The failure of cutaneous integration directs epithelial cells to migrate or grow down along the implant's surface during the final stage of wound healing, resulting in a phenomenon known as “marsupialization” or epidermal downgrowth [29]. As shown in Figure 1, this complex healing process leads to distal sinus tract formation. At the same time, FBR directs the formation of fibrous capsules proximally [32, 33, 34, 35]. The sinus tract that forms around the distal part of the implants is always colonized by bacteria but also can serve as a breeding ground for opportunistic pathogens as well as commensal microbes [29]. When a potential pathogen reaches a critical number, it can lead to a clinically significant infection. These are reported to range from 0% to 56% [36, 37]. Thus, the need for further research studies to find a way to establish a stable biological seal at the stoma and provide a long‐term solution for patients wanting to undergo this treatment.

Figure 1.

A schematic illustration of epithelial downgrowth along a percutaneous implant surface. The epidermal cells migrate downward along the implant surface, encasing it circumferentially without establishing direct integration, i.e., biological attachment.

In an attempt to achieve stable integration of the epidermis with the implant surface, Pendergrass et al. [38] investigated the use of hydroxyapatite (HA)—the primary inorganic component found in antlers—as a coating on titanium implants [38, 39, 40]. Prior to this, Jensen and his colleagues also explored HA as a percutaneous coating material in translational animal models [41, 42, 43]. However, data from both groups demonstrated the limited utility of HA at longer‐term follow‐up, and HA‐coated titanium implants failed to improve cutaneous integration in clinical use [41, 42, 43, 44]. There are three possible reasons for these results. First, antlers protrude through the deer's scalp only during the rutting season, after the velvet is shed [40], implying that the skin seal is required only for short‐term protection. Second, HA is naturally resorbed and reused within the body by macrophages and osteoclasts [45, 46, 47, 48], which may lead to the rapid degradation of the HA coating. Third, HA coating methods and the resulting crystallinity significantly influence its resorption rate [49, 50, 51]. With time, these coatings could be resorbed, exposing the base Ti surface, resulting in loss of stable seal. Despite the apparent failure of HA coatings to achieve persisting cutaneous integration, HA‐coated bone‐anchored hearing aids (BAHA) appear to work [52]. Clinical studies in these cases suggest that HA‐coated abutments produce a thick layer of skin, just like the junctional epidermis [52]. An additional report of HA‐coated titanium implants in the craniofacial skeleton, showed that this approach can work in the head and neck region. Although limited, these data suggest that the use of HA coatings on titanium implants may warrant further investigation [53].

Previous studies also may have overlooked a critical apatite‐based percutaneous ectodermal organ: the tooth. It is well established that the junctional epithelium attaches to tooth enamel via specialized cell junctions, hemidesmosomes (HDs), ensuring the stability of the junctional gingival seal with the tooth surface for its lifetime [54, 55]. It has been suggested that enamel surfaces induce a unique cellular response in gingival mucosal cells of the junctional epithelium, leading to HD expression [56, 57, 58]. Such evidence suggests that certain bioactive apatites could induce appropriate cellular responses, allowing epithelial cells to adhere to non‐living surfaces. A crucial factor is that tooth enamel is composed mainly of fluoridated apatite, known as fluorohydroxyapatite (FHA) or fluorapatite (FA), which shows greater crystallinity when compared to HA [59, 60, 61]. Most importantly, its ability to resist bioresorption over the lifetime of the animal suggests that it has significantly lower resorption rates when compared to HA. Based on these facts, our group has investigated FA as an interface material for percutaneous device applications, with the rationale of mimicking the composition and crystallinity of the enamel surface and creating a “dental‐like” structure. Our previous work demonstrated that FA, when sintered at high temperatures (1050°C–1250°C), increased the activity of fibroblasts and keratinocytes in vitro while maintaining the high crystallinity in the FA structure [62]. Most strikingly, small animal (rat) feasibility data showed a reduction in GT and epidermal downgrowth at the skin‐device interface 12 weeks post‐implantation [63]. Histomorphometry and molecular analyses revealed that GT surrounding FA‐coated implants closely resembled healthy dermal tissue when compared to uncoated titanium implants, suggesting the ability of FA to promote a normal healing process at the implant exit site [63].

Many translational animal models—rats, guinea pigs, rabbits, and pigs—have been used to study the effects of healing around biomimetically coated percutaneous titanium implants [63, 64, 65, 66]. However, none of these in vivo models involved weight‐bearing. During daily ambulation, the periprosthetic skin also endures constant tensional, shear, and torsional forces, creating a more mechanically challenging interface environment than the non‐weight‐bearing interfaces tested in the previously mentioned models. Since FA coating has successfully demonstrated its ability to reduce downgrowth and improve healing cascades in the rat model [63], the current study aimed to confirm this finding by using our previously reported sheep amputation and load‐bearing implantation model [27, 36, 67, 68, 69, 70, 71]. We hypothesized that coating percutaneous titanium OI devices with an FA coating would reduce epidermal downgrowth and facilitate GT healing at the skin‐device interface, thereby improving stomal wound healing and, by doing so, provide a biologically stable stoma.

2. Methods

2.1. Material and Chemicals

Unless otherwise specified, all reagent‐grade chemicals were purchased from Sigma Aldrich (St. Louis, MO).

2.2. Implant Fabrication and Coating

A previously validated intramedullary implant, specifically designed to fit the fused sheep metacarpal III/IV bone, was utilized in this study [27, 36, 67, 68, 69, 70, 71]. Detailed specifications of the implants have been published elsewhere [27, 68]. In short, the original engineering schematics were used to manufacture the percutaneous OI implants, which comprise both endo‐ and exo‐prosthetic components (Figure 2). These components were fabricated using medical‐grade titanium alloy (Ti6Al4V) and grit‐blasted at the proximal endo‐prosthetic end (DJO, Austin, TX). The distal end was coated with a porous titanium coating (Thortex, Portland, OR) to facilitate bone ingrowth and promote skin interdigitation for preventing shear forces at the interface.

Figure 2.

A photograph showing a fabricated percutaneous osseointegrated implant for the sheep fused 3–4 metacarpal bone attached to the exo‐prosthesis (foot). The entire endo‐prosthetic and percutaneous portions of the implant were coated with FA.

To evaluate the efficacy of FA as a coating, the entire implant was micro‐blasted with FA powder, which was pre‐sintered at 1150°C and sieved to achieve particle sizes in the range of 25 < d < 75 µm. For depositing a thin coating, a microblaster (AccuFlo Microblaster; Comco Inc., Burbank, CA) was used with an attached 0.046” nozzle after optimizing the nozzle size for uniform FA deposition. Moreover, the nozzle was positioned 2 cm from the implant surface. Each implant was uniformly coated using a 10‐min spray with a constant nitrogen gas flow of 65 PSI. The coated surfaces were subsequently characterized using scanning electron microscopy (SEM; FEI Quanta 600F, Hillsboro, OR) coupled with energy‐dispersive spectroscopy (EDS; EDAX, Pleasanton, CA) to confirm the coating uniformity, and x‐ray photon spectroscopy (XPS; Kratos Axis Ultra DLD, San Diego, CA) was used to confirm the elemental composition of the coated surfaces.

2.3. Animal Study Design

Five skeletally mature Rambouillet sheep (n = 5, > 18 months old) underwent amputation followed by implantation of an OI FA‐coated implant, as described below, using an institutionally approved IACUC protocol (University of Utah 1721). A historically successful surgical protocol was replicated to minimize the numbers of animals used and to serve as the control group (n = 5, uncoated Ti) [31]. Following surgery, animals were monitored for 12 weeks. At necropsy, the implants and surrounding tissues were collected for analysis of epithelial downgrowth and GT analyses and compared to historical controls. Healing at the implant exit site was assessed using standard histology and immunohistochemistry (IHC) techniques. To limit surgical variability, all procedures were performed by a single orthopedic surgeon.

2.4. Surgical Procedure

The single‐stage surgical procedure followed a previously published surgical protocol [27, 36, 71]. Briefly, animals were fasted overnight, and transdermal fentanyl patches (1–2 µg/kg/h) were applied 24 h before surgery and maintained for a further 72 h after surgery. Water was withheld on the morning of the procedure. General anesthesia was induced with propofol (4–10 mg/kg/IV) and maintained with isoflurane (0.5%–5% in oxygen). Mirroring previous studies, the right forelimb was shaved, cleansed with 3‐time alternating scrubs of Betadine and 70% alcohol, and covered with sterile drapes for surgery. A temporary tourniquet was placed proximal to the carpal‐metacarpal joint. Following exposure, the flexor and extensor tendons were tenodesed in a neutral position using 0 FiberWire suture (Arthrex, Naples, FL), and the amputation was then performed by transecting the fused metacarpal 3–4 bone at the metaphyseal flare. Post amputation, the medullary canal was reamed, using frequent saline flushes to avoid heat‐induced bone necrosis. An appropriately‐sized FA‐coated implant from a selection of five sizes was press‐fitted into the bone using a hand mallet. The implant's Morse‐tapered post was routed through a stab incision, preserving the anterior blood supply and venous drainage of the skin flap. The skin incision was closed using 3‐0 Vicryl sutures in layers, leaving the corners open for drainage. An exo‐prosthetic hoof was attached to the implant and gently impacted to secure it. Post‐surgical radiographs were taken to confirm proper alignment (Figure 3). The wound was dressed with Telfa pads and vet wrap, with dressing changes weekly for the first week and biweekly thereafter.

Figure 3.

A representative post‐implantation radiograph showing the implant positioned within the sheep's fused 3‐4 right metacarpal bone.

2.5. Postoperative Treatment and Monitoring

Postoperative treatment included Excede (6.6 mg/kg) preoperatively and daily for 5–7 days to prevent infection and carprofen (4.4 mg/kg) once a day for a miniumum of 7 days for pain management, with additional doses prescribed as needed. A fentanyl trans‐dermal patch was applied the day prior for surgery, and then replaced every 72 h for at least 216 h. Animals were observed twice daily for 2 weeks, then three times per week, to monitor general health, skin flap health, and weight‐bearing status. One animal was excluded at Week 4 due to implant failure resulting from accidental trapping of the hoof under the steel pen enclosure. Also, the prosthetic hoof was changed as needed. Animals were sacrificed at 12 weeks post‐surgery using the approved procedure, which followed the American Veterinary Medical Association guidelines.

During necropsy, the right forelimb was shaved and disarticulated at the carpal‐metacarpal joint. Post‐harvest, two sections of the interfacial tissues at the implant exit sites were collected from the anterior/posterior side for molecular analyses. For samples intended for IHC, collected tissues were marked and embedded in optimal cutting temperature (OCT) compound (Thermo Fisher, Waltham, MA) in a plastic mold and frozen using an isopentane bath pre‐chilled over dry ice. The frozen OCT‐embedded tissue was wrapped in aluminum foil and stored at ‐80°C until sectioning. The rest of the harvested limb with the implant still intact with the bone was processed for histological analysis (see below).

2.6. Histological Processing and Analyses

To analyze the skin‐device interface, tissue samples were fixed in 10% neutral formaldehyde and then dehydrated in ascending grades of alcohol using an automated tissue processor (Tissue Tek VIP; Sakura Finetek, Torrance, CA). They were then embedded in poly (methyl methacrylate) (PMMA). These PMMA blocks were sectioned using a precision saw (Isomet 4000; Buehler, Lake Bluff, IL) to obtain 2 mm thick sections, glued to a plastic slide, and then ground and polished to achieve thin, optically finished, smooth sections using an automated grinder polisher (Ecomet 300; Buehler, Lake Bluff, IL).

Histological evaluation of the skin interface was conducted post‐staining with hematoxylin and eosin (H&E) using a previously published method [72]. Briefly, the slides were immersed in a heated Wiegert hematoxylin mixture (at 60°C) for 10 min, followed by bluing under running tap water. Subsequently, the slides were counterstained with 1% eosin for 2 min at 60°C. Finally, slides were cleaned with 70% ethanol, dried, and examined under a light microscope (Nikon, Melville, NY). Mirroring the previously published methodologies [63, 73], the downgrowth and GT area were measured using the Nikon Elements software (NIS; Nikon, Melville, NY) with the line measurement and area measurement tools, respectively. Four measurements were made from each of four animals, resulting in 16 measurements per group for both downgrowth and GT area.

2.7. Immunohistochemistry

To assess the expression of selected wound‐healing‐related inflammation markers, fresh‐frozen periprosthetic tissues were used. Ten‐micrometer‐thick sections were obtained using a cryotome (Thermo Fisher Scientific, Waltham, MA), and these sections were transferred onto charged microscope slides and stored at −80°C until further use. Prior to staining, the samples were fixed in 100% acetone for 15 min at room temperature, followed by a 5‐min wash in PBS; this fixation process was repeated three times. The slides were allowed to dry at room temperature for 15 min and then permeabilized with 0.3% Triton X‐100 in PBS for 1 h.

To minimize non‐specific binding, samples were initially blocked using BlockAid blocking solution (Thermo Fisher Scientific, Waltham, MA) for 1 h at room temperature. Following blocking, the slides were incubated overnight at 4°C with primary antibodies targeting transforming growth factor α (TGF‐α; HPA042297, Sigma‐Aldrich, St. Louis, MO), epidermal growth factor receptor (EGFr; ab52894, Abcam, Waltham, MA), collagen IV (COLIV; ab6586, Abcam, Waltham, MA), or keratin 6 (KRT6; HPA061168, Sigma‐Aldrich, St. Louis, MO). The following day, samples were incubated with the corresponding secondary antibodies (Abcam, Waltham, MA) for 1 h at room temperature. Antibodies targeting F‐actin (ab176752, Abcam, Waltham, MA) were also added to the samples containing TGFα and EGFr and incubated for 1 h at room temperature. After incubation, slides were washed thrice for 5 min each in PBS and mounted with a DAPI‐containing mounting medium (ab104139, Abcam, Waltham, MA). Spatial expression of TGFα and EGFr were captured using a confocal microscope (Olympus FV1000, Olympus, Japan), and percent fluorescence signals were quantified using ImageJ (NIH, Bethesda, MD) [74].

2.8. Statistical Analysis

All data is reported as means ± SEM. Statistical differences in downgrowth and GT area were calculated using mixed effects linear regression models, as these were clustered datasets with four repeated measurements per animal. This was used to account for the lack of independence that could be introduced by data clustering. A p‐value less than 0.05 was considered significant. All statistical analyses were conducted using STATA (StataCorp, College Station, TX).

3. Results

The XRD patterns of the in‐house synthesized FA closely matched the reference (RRUFF R060421). Crystallinity was calculated to be approximately 94.5% for samples sintered at 1150°C (Figure 4).

Figure 4.

Representative XRD patterns of FA powder sintered at 1150°C, which was used for low‐temperature coating, and a fluorapatite reference (RRUFF R060421).

After coating the disks with FA using the microblaster, the coated implant surfaces were analyzed using EDS to assess the uniformity of the deposition. EDS imaging confirmed the presence of FA on the titanium surface (Figure 5). Semi‐quantitative EDS analysis revealed atomic weight percentages of calcium (Ca), phosphorus (P), and fluorine (F) to be 22.1%, 12.6%, and 3.4%, respectively (Table 1). In total, the data accounted for only 3.4% of the uncoated surface (elements representing Ti alloy). As expected, the surface of the uncoated control implant showed only Ti, aluminum, vanadium, and oxygen (from the oxidized titanium oxides). It is worth noting that while adsorbed oxygen made up 53.2% of the surface composition of coated surfaces, only 3.8% was found in the uncoated Ti surface.

Figure 5.

(A and B) A set of scanning electron micrographs depicting the surface of the Ti substrate with no coating. (C) A representative EDS elemental map of the cross‐section of uncoated Ti6AlV4. (D and E) A representative set of scanning electron micrographs showing FA‐coated Ti substrate. (F) A representative EDS elemental map of the FA‐coated cross‐section. The EDS spectrum analysis revealed the elemental composition of the coating with prominent peaks corresponding to F, Ca, and P, confirming the presence of a calcium phosphate‐rich layer on the surface. Also, uncoated base substrate, peaks representing Ti‐6Al‐4V (i.e., Ti, Al, and V peaks) were also present post‐coating, indicating less than 100% overage of the surface with FA coating.

Table 1.

Relative surface elemental compositions of uncoated and FA‐coated titanium devices as determined by EDS data.

	Elements	F	O	Ti	Ca	P	Al	V
Uncoated Ti	Atomic %	—	3.8	80.3	—	—	12.3	3.6
FA‐coated Ti	Atomic %	3.4	53.2	3.4	22.1	12.6	5.1	0.2

Additionally, XPS analyses were conducted to confirm the relative atomic percentage of the FA coating on the Ti alloy surfaces. Figure 6 displays a representative scan with the elements labeled on the peaks. Table 2 presents the corresponding relative atomic percentages of the components.

Figure 6.

Top: A representative XPS spectrum of the O 1s, Ti 2p, Ca 2s, Ca 2p, C 1s, Al 2s, and Al 2p regions for the uncoated titanium alloy surface. Bottom: A representative XPS spectrum of the F 1s, Ca 2s, Ca 2p, P 2s, P 2p, C 1s, and O 1s regions for the FA‐coated titanium alloy surface. The presence of distinct peaks at ~684 eV (F 1s), ~347 eV (Ca 2p₃/₂), and ~133 eV (P 2p) confirms successful deposition of FA. The O 1s spectrum shows contributions from both phosphate (PO₄³⁻) and TiO₂. Ti and Al 2p signals are significantly attenuated, indicating a continuous and thin coating.

Table 2.

Relative surface elemental compositions of uncoated and FA‐coated titanium devices calculated from XPS data.

Elements	F	O	Ti	Ca	C	P	Al
Uncoated Ti (Atomic %)	—	31.0	9.3	1.2	55.6	—	1.2
FA‐coated Ti (Atomic %)	3.9	48.3	0.5	16.5	17.2	12.8	0.7

As stated in the methods section, five female Rambouillet sheep underwent amputation surgery and were implanted with a single FA‐coated device in the right metacarpal. One sheep was removed from the study due to injury. The remaining sheep survived to the endpoint without any signs of infection or inflammation at the implant exit site. At necropsy, skin samples were collected for IHC, and skin‐implant samples were collected for H&E staining.

Histomorphological evaluations were conducted on PMMA‐embedded and H&E‐stained sections. The data of this study group were compared to historical controls to assess downgrowth and GT area, with each group having a total of 16 (n = 16) measurements. While the FA‐coated implant group exhibited little to no downgrowth from the position at the time of surgery to its position at 12 weeks post‐healing, the uncoated Ti control group had noticeable downgrowth (Figure 7). In the FA‐coated group, when compared to the control group, only small areas of GTs were noticeable around the post at the implant exit site.

Figure 7.

A representative set of H&E‐stained cross sections of the tissue at the skin‐implant interface. Top row: Uncoated control samples. Bottom row: FA‐coated samples showing the granulation tissue (GT) area (blue circles) and downgrowth (blue arrows) at 12 weeks post‐surgery. The last three right‐hand images represent increasingly magnified implant exit sites (blue circles), showing morphologies of periprosthetic tissue. Effectively, fiber‐dense tissue morphology was present within the FA‐coated groups. In contrast, a cellular dense GT was present in the control group. Implant exit sites are indicated by white arrows pointing toward the bottom of the images. It is important to note that the area of GT is decreased (if present) in samples with FA‐coated implants compared to uncoated Ti implants.

Statistical analyses of GT areas revealed a significant decrease (p < 0.001) in the FA‐coated group compared to the uncoated control (Figure 8). Similarly, epithelial downgrowth was also significantly reduced (p < 0.001) in the FA‐coated device group relative to the control (Figure 8).

Figure 8.

Left: A bar chart illustrating the average granulation tissue area around the device at 12 weeks after implantation. The FA‐coated implants exhibited a significantly (p < 0.001) reduced GT compared to the uncoated Ti implant. Right: A bar chart presenting the average downgrowth from the position at surgery to the position after 12 weeks post‐implantation. The data indicate a significantly (p < 0.001) decreased downgrowth in FA‐coated implants relative to uncoated Ti implants.

To determine whether wound healing protein markers were expressed in tissue samples, IHC studies were undertaken. For this analysis, EGFr or TGFα were selected as surrogate markers due to their role in wound healing, and as they were differentially expressed in our previous studies [55, 63, 66, 75]. Compared to the uncoated control groups, the FA‐coated implant groups exhibited a decreased presence of fluorescent signals for both EGFr (Figure 9) and TGFα (Figure 10) within the GT. Additionally, a decrease in the migration marker KRT6 is seen in the epidermis of the FA‐coated implant group (Figure 11).

Figure 9.

A representative set of confocal micrographs showing the presence of EGFr within the granulation tissue at the interface of the skin and the implant. Green fluorescence indicates actin, while pink indicates EGFr expressions. There was more EGFr signal present in the granulation tissue area on the uncoated control compared to the FA‐coated implant.

Figure 10.

A representative set of confocal micrographs showing the presence of TGFα in the granulation tissue at the interface of the skin and the implant. Green is indicative of β‐actin, and pink is indicative of TGFα. There was more signal present in the granulation tissue area on the uncoated control compared to the FA‐coated implant.

Figure 11.

A representative set of confocal micrographs showing the presence of KRT6 in the granulation tissue at the interface of the skin and the implant. Red is indicative of KRT6, and blue is DAPI staining for the nucleus. There was more signal present in the epidermis on the uncoated control compared to the FA‐coated implant.

4. Discussion

This study was designed to evaluate the efficacy of FA‐coated percutaneous OI implants to prevent or reduce epithelial downgrowth under weight‐bearing conditions, leveraging the known ability of FA to promote periprosthetic epidermal cell adhesion. Our findings—including clinical and histological observations, histomorphometric data (downgrowth length and GT area), and IHC assessments—supported this hypothesis. The H&E‐stained sections revealed a marked reduction in cellular infiltration of the GT at the implant exit site, as well as a reduction in epithelial downgrowth in the FA‐coated group compared with controls (Figure 7). Complementary IHC analysis also demonstrated downregulation of wound‐healing markers [76], such as EGFr and TGFα, suggesting a stable and mature epithelial interface (Figures 9 and 10), as well as the absence of keratin 6 (KRT6) at the interface (Figure 11). Notably, there were statistically significant reductions (p < 0.05) in epithelial downgrowth and GT areas at the implant exit sites of FA‐coated devices (Figure 8), supported by our previous finding in rats [63].

It should be emphasized that the healing outcomes data generated in this research were indirect measures of the FA coating's ability to form HD‐based adhesion with the terminal epithelial cells to prevent continuous, ongoing wound healing, as commonly observed in all percutaneous devices. Visualization of HD at the interface using transmission electron microscopy (TEM) would have provided direct evidence of HD plaques between epithelial cells and implant surfaces. However, the embedding technique used in this study was unsuitable for directly visualizing HD plaque structures. Future studies should, therefore, employ definitive methods capable of directly confirming HD formation to further substantiate this mechanism of cutaneous integration.

The use of apatite coatings to reduce epithelial downgrowth and promote healing around percutaneous OI implants is not a novel concept. HA, in particular, has been utilized in intraosseous transcutaneous amputation prostheses, various dental implants, and BAHAs with mixed clinical outcomes [8, 10, 39, 77, 78, 79]. In these studies, the HA coatings have been shown to promote a soft tissue seal at the skin‐device interface during short‐term translation studies, but they have demonstrated limited clinical success in human patients [8, 52, 80]. Among these clinical devices, BAHA has shown the most effectiveness. Most likely, this is because BAHA devices are non‐weight‐bearing and are placed in an area of the skull where the soft tissues can be easily immobilized by thinning and adhesion onto the peri‐implant periosteum [81].

In contrast, both intraosseous transcutaneous amputation prostheses and dental applications of HA exhibit premature resorption of the HA coating, thereby reducing the effectiveness of the approach [82]. When investigating the HA coating techniques, the most common method for HA deposition for clinical application is plasma spraying, which involves heating the apatite to extremely high temperatures (up to 16,000°C) [83]. At this temperature, molten HA particles are propelled onto the implant surface via a plasma jet and rapidly cooled to form a coating [84, 85]. Although plasma spray creates a uniform HA coating on the surface, the high heat required for this process ultimately reduces the apatite coating's crystallinity to around 65%–70% [86, 87], and probably a proportion of the coating may phase transform into other Ca‐P forms, such as beta‐tricalcium phosphate [88, 89, 90]. Although both HA and beta‐tricalcium phosphate are bioactive and osteoconductive [91], and suitable for orthopedic applications [90], the current percutaneous application requires high crystallinity, similar to that of dental enamel, where the longevity of the coating is of paramount importance [92, 93, 94].

In comparison, a highly crystalline FA form is sparingly biosorbable, just like human tooth enamel, which provides long‐term durability and stability. Its presence at the tissue‐implant interface creates a durable protective layer and promotes the formation of HDs with the junctional epithelium, resulting in a robust, lifelong soft‐tissue seal [54, 55]. This is particularly critical in weight‐bearing applications, where mechanical forces at the skin‐implant interface can compromise healing and integration [95, 96]. Thus, in this study, we combined the use of FA with high crystallinity (94.5%, Figure 4) and used a low‐temperature microblasting technique to address two significant foreseen limitations of plasma‐sprayed HA coatings for percutaneous application; these are increased bioabsorbability and thermal degradation, respectively. This strategy appeared to provide a more stable coating, at least for the duration of this 12‐week study.

As stated, this study employed a low‐temperature microblasting coating technique to preserve the high crystallinity of sintered FA powders. Both EDS and XPS (Figures 3 and 4; Tables 1 and 2) corroborated the presence of a thin FA coating on the implant substrate. The data also confirmed the utility of the above surface surveying techniques in apatite coating applications. The XPS scans of the coated surfaces revealed distinct peaks at approximately 684 eV (F 1s), 347 eV (Ca 2p₃/₂), and 133 eV (P 2p), verifying the presence of FA. In addition, attenuation of Ti and Al 2p signals relative to uncoated controls indicated effective coverage of the FA layer on the titanium alloy surface. The SEM images and corresponding EDS analyses further confirmed the uniformity of the coating. While ~81% of the uncoated surfaces consisted of Ti, only 3.4% Ti was detected on the coated samples, signifying substantial surface coverage of the employed coating technique. It is worth noting that XPS only probes the uppermost 2–5 nm of the surface, whereas EDS can penetrate to depths of approximately 0.1–3 µm [97, 98], suggesting that the FA coating thickness likely falls within the micron range, as both techniques confirmed coating coverage. Nonetheless, additional techniques are needed to precisely quantify the coating thickness. One could employ profilometry, ellipsometry, X‐ray microanalysis, and cross‐sectional SEM for these analyses [99, 100]. Although various powder coating techniques, such as plasma spray and sputter coating, exist for depositing apatite materials onto titanium surfaces at a 1 mm scale, they also reduce the crystallinity of FA and increase the byproducts of FA due to the high temperature, which could aid the faster resorption of the coating [83, 85, 92, 101]. As the current macroblast technique is not capable of depositing a thick coating, comparative investigations of coating techniques are warranted to optimize coating techniques and performance in future studies.

The most notable finding was the statistically significant differences in GT tissue area at the implant exit sites between the groups (p < 0.001). The increased area of GT with high cellular density observed in the uncoated devices suggested a more pro‐inflammatory microenvironment, potentially resulting from immune responses to bacterial infiltration, which is facilitated by the lack of an epithelial seal at the skin‐implant interface. In contrast, the coated devices appeared to promote more favorable healing outcomes, with reduced cellular infiltration within the GTs, suggesting less inflammatory, pro‐healing environments, as evidenced by reduced EGFr and TGFα expression, which further supported our hypothesis. The EGFr is known to play multiple roles during the early phases of wound healing [102, 103], including promoting inflammation, angiogenesis, and re‐epithelialization [102]. Their presence is enhanced by neutrophil infiltration [102, 104]. Thus, the reduced EGFr expression on coated surfaces could indicate successful epidermal integration with the implant surface. This could also limit bacterial ingress and associated immune responses within the periprosthetic soft tissues (Figure 9). Additionally, EGFr is involved in stimulating epithelial cell proliferation and migration [105], further emphasizing their role in wound dynamics of periprosthetic tissue at the implant exit site.

It is worth noting that reduced expression of TGFα (a part of the EGF superfamily) was noted in the coated group, perhaps a better indicator of the coating's ability to limit downgrowth (Figure 10), as TGFα is known to stimulate keratinocyte migration and proliferation [106]. It is widely accepted that, during the re‐epithelialization stage of the normal wound‐healing process, the avascular epidermis migrates along the highly vascular GT, seeking signals to form tight cell‐to‐cell connections and reestablish its barrier function [107, 108]. When this cell‐to‐cell connection is formed, the so‐called migrating epidermis expressing KRT6 can then switch phenotypes and start differentiating to the upper layers of epithelium expressing KRT1 and KRT10 [109, 110]. This marks the end of the re‐epithelialization phase of wound healing, and concludes the initial wound‐healing process. As shown in numerous studies of percutaneous OI devices, the presence of skin‐protruding implants disrupts the typical wound‐healing trajectory [63, 111, 112, 113, 114]. In addition, the presence of an area of vascularized periprosthetic GT instigates epithelial migration along the implant surface, resulting in clinically observed downgrowth [29]. It is important to note the absence of KRT6 on the coated devices, when compared to the control titanium surface (Figure 11). The data from this study further support the notion that (similar to dental enamel), establishing a HD connection between the implant surface and the epithelial cells may prevent the exacerbated downgrowth commonly observed with percutaneous devices [55]. This conclusion is supported by the statistically significant differences in downgrowth observed between the coated and uncoated groups.

Interestingly, TGFα was upregulated in a previous rat study [63]. In contrast, in this study, its expression was downregulated. Although puzzling, it may be related to the stages of wound healing themselves. In the previous rat study, rats were sacrificed 4 weeks post‐implantation, but in this study, sheep were observed for 12 weeks. Literature indicates that TGFα expression level in non‐ischemic wounds varies during wound healing [115]. Also, there may be species differences in wound healing between rats and sheep. Therefore, it is reasonable to conclude that, at the completion of wound healing with FA‐coated devices, expression of TGFα expression is expected to subside. Moreover, both EGF and TGF chemokines are expressed during the re‐epithelialization phase of wound healing [102]. Since epithelial migration and the subsequent downgrowth are ongoing when no coating was used (i.e., control), the presence of EGF and TGF was expected and has been supported as seen in Figures 6 and 7. It should be pointed out that EGF and TGFα share receptor EGFr, and when bound to EGFr, enhance cell proliferation to promote wound healing [76, 116, 117].

This study was a proof‐of‐concept study demonstrating the importance of surfaces in regulating and promoting wound healing around percutaneous devices. Two major limitations exist with this study; they are: the small sample sizes for each group and the short follow‐up. Increasing the number of animals per group would be beneficial to confirm the presence of any differences between groups. Increasing the length of follow‐up would help to confirm the stability of any interaction between the epithelium and the FA‐coating—an important prerequisite for any future human studies. In addition, historical samples were used for the uncoated Ti control. The long‐term mechanical and chemical stability of the coating was not studied, and a degradation study of the coating on the surface, as well as the coating thickness, needs to be conducted. Furthermore, this study focuses on only one type of biomaterial coating and one type of coating technique. Future studies should include similar bioactive biomaterials, time series, and bulk RNA sequencing, as well as single‐cell RNA studies, to investigate the underlying mechanisms. Since clinical studies indicate that cutaneous OI devices have an infection rate of up to 56% [36, 37], we also need to test rates of infection in an effective infection model. Finally, although we suspect that the FA played a role in promoting expression of the HDs at the skin‐device interface, thereby limiting downgrowth around the stoma, no studies were performed to quantify the level of HD expression. Such quantification is extremely difficult in the translational model and should be conducted in cell‐culture studies, which are currently being undertaken.

5. Conclusion

In a well‐established sheep amputation model, FA‐coated weight‐bearing percutaneous OI devices demonstrated reduced epithelial downgrowth and smaller regions of GT with markedly reduced cellular infiltration compared to uncoated titanium controls. IHC analysis revealed significant differential expressions of EGFr and TGFα, further supporting the role of FA coating in modulating wound healing responses at the soft tissue–implant interface. Collectively, these findings indicate that FA coating represents a promising strategy for promoting stable healing at percutaneous implant exit sites.

Author Contributions

Samantha Steyl contributed to animal studies, data acquisition, documentation, data analysis and interpretation, and manuscript drafting and critical review of the submission. James Peter Beck and Jay Agarwal contributed to the study design, performed animal surgeries, and critically reviewed the manuscript. Jill Shea contributed to the study design, data acquisition and analysis, drafting, and critically reviewed the manuscript. Ruben Sundramurti contributed to the acquisition, analysis, and interpretation of histological data and critically reviewed the final manuscript. David Rou contributed to the acquisition, analysis, and interpretation of immunohistochemistry data and critically reviewed the final submission. Kent N. Bachus helped to acquire funding for the study, contributed to the study design, and provided a critical review of the manuscript. Sujee Jeyapalina acquired funding for the study and contributed to the study design, data acquisition and analysis, drafting, critical review, and final submission of the manuscript.

Data Availability Statement

The data that support the findings of this study are available from the corresponding author upon reasonable request.